Biphasic implant device transmitting mechanical stimulus

ABSTRACT

Tissue implants prepared for the repair of tissues, especially avascular tissues such as cartilage. One embodiment presents an electric potential capable of receiving and accumulating desirable factors or molecules from surrounding fluid when exposed to dynamic loading. In another embodiment the implant promotes tissue conduction by retarding, restricting and controlling cellular invasion through use of gradients until competent tissue forms. Further embodiments of the tissue implants may be formed into a multi-phasic device that provides deep tissue mechanical stimulus by conduction of mechanical and fluid forces experienced at the surface of the implant.

BACKGROUND OF THE INVENTION

What is disclosed is a device for repairing and replacing lost ordamaged tissue. Particularly, one embodiment is directed to amulti-phasic prosthetic device for repairing or replacing cartilage orcartilage-like tissues. Said prosthetic devices are useful as articularcartilage substitution material and as a scaffold for regeneration ofarticular cartilaginous tissues.

Cartilage is found throughout the body, such as in the supportingstructure of the nose, ears, ribs (elastic cartilage), within themeniscus (fibrous cartilage), and on the surfaces of joints (hyalinecartilage or articular cartilage). A joint is a bending point where twobones meet. The knee, hip, and shoulder are the three largest joints.

The specialized covering on the ends of bones that meet to form anarticulating joint is called hyaline or articular cartilage. It is thecartilage that is damaged and wears as one ages, or sustains an injury.Articular cartilage is unique amongst the body tissues in that it has nonerves or blood supply. This means that damage will not be felt untilthe covering wears down to bare underlying bone. Bone is very sensitiveand the sharp pain of arthritis often comes from irritation of bonenerve endings and since human tissue has a very limited capacity to healwithout a blood supply, articular cartilage cannot repair itselfeffectively.

Articular cartilage tissue covers the ends of all bones that formdiarthrodial joints. The resilient tissues provide the importantcharacteristic of friction, lubrication, and wear in a joint.Furthermore, it acts as a shock absorber, distributing the load to thebones below. Without articular cartilage, stress and friction wouldoccur to the extent that the joint would not permit motion. As statedabove, articular cartilage has only a very limited capacity toregenerate. If this tissue is damaged or lost by traumatic events, or bychronic and progressive degeneration, it usually leads to painfularthrosis and decreased range of joint motion.

Articular cartilage repair following injury or degeneration represents amajor clinical problem, with treatment modalities being limited andjoint replacement being regarded as appropriate only for the olderpatient.

Current treatments for articular cartilage damage are varied and includeanti-inflammatory medication, viscosupplementation, arthroscopicchondroplasty, autogenous articular cell implantation, microfracture andosteochondral articular transplantation.

Anti-inflammatory medication: Aspirin was the first anti-inflammatorymedication in the world. This was followed in 1950 by cortisone(steroidal medication) used orally or by injection. (Extensive use ofcortisone not only has a wide variety of harmful effects, but is alsobelieved to harm cartilage.) Later the non-steroidal drugs such asMotrin came along. These were safer than Aspirin and cortisone but hadpotent side effects, especially causing bleeding within the stomach andintestinal ulcers. These complications led to the development of theCOX-2 inhibitor drugs, Celebrex and Vioxx. While much safer andseemingly more effective, Vioxx was found to have significant cardiacside effects and is no longer available. With certain precautions,Celebrex is still widely used. However, these anti-inflammatorymedications only treat the symptoms of cartilage damage and arthritisand do not promote repair.

Viscosupplementation: Viscosupplementation is a procedure that involvesthe injection of gel-like substances (hyaluronates) into a joint tosupplement the viscous properties of synoval fluid. Currently,hyaluronate injections are approved for the treatments of osteoarthritisof the knee in those who have failed to respond to more conservativetherapy, Once again, this procedure only treats the symptoms ofcartilage damage and arthritis and does not promote repair.

Arthroscopic chondroplasty: Chondroplasty is a term referring to thearthroscopic smoothing of unstable articular surfaces either withmechanical shaving or thermal devices. While not a restorative measure,so called debridement can be useful in reducing irritating cartilagedebris that breaks off in the joint or causes catching or grindingsensations. The resulting improvement in the control of inflammation canlast for several years. But this is not a final solution as thedegenerative process continues to wear away at the articular cartilage.

Autogenous articular cell implantation (ACI): Autogenous cellimplantation can be used for large, shallow defects, which do notinvolve the subchondral bone. In this procedure, cartilage cellscollected from the patient and grown to many millions through cellculture techniques are injected into the joint, under a membrane thathas been attached to the cartilage surface. Although successful, thewindow of opportunity for this procedure is often missed, as the fewclinical symptoms showing the need for this treatment are not evidentuntil the defect deepens to involve the underlying bone, thus the damageencountered upon detection is frequently too extensive for repairthrough ACI.

Microfracture: The goal of this arthroscopic technique is to improve theblood supply to the bare areas of the joint by creating tinyperforations in the underlying bone. The resulting bone marrow bleedingcarries powerful growth stimulating factors found in platelets as wellas stem cells to the damaged area creating what is referred to as asuper-clot. Healing and repair follow over several weeks. Studies haveshown that microfracture techniques do not fill in the chondral defectfully and the repair material formed is fibrocartilage. Thefibrocartilage tissue can temporarily return function for activitiessuch as running and a sport play, but ultimately fails, asfibrocartilage is unable to mechanically share and disipate loadingforces as effectively as the original hyaline cartilage. Fibrocartilageis much denser and isn't able to withstand the demands of everydayactivities as well as hyaline cartilage and is therefore at higher riskof breaking down.

Osteochondral articular transplantation: Osteochondral transplantation(i.e. mosaioplasty) involves transportation of tissue plugs from onelocation of the knee to another. Special instrumentation has beendevised to harvest plugs of articular cartilage and its supporting bonefrom the patient's own joint. The harvested tissue is then transportedto the damaged site where it is inserted into prepared holes. Severalplugs can fill up rather larger defects and will grow to re-supply a newjoint surface. Unfortunately, this procedure leaves defects of equal orworse proportions elsewhere and often the harvested tissue is not viabledue to the traumatic harvesting procedure.

Due to the problems associated with current state of the art treatments,much work has been done to produce a synthetic off-the-shelf scaffold tobe used in place of the harvested osteochondral plug.

Originally, single-phase scaffolds of uniform construction werecontemplated for use as implants. However, these single-phase scaffoldimplants proved unsuccessful in healing of the complex multiphasicarticular cartilage along with the underlying bone. Soon biphasic andthen gradient devices were developed that were either mechanically oranatomically specific for the tissues involved. While these showed animprovement over single phase devices, it is evident that these devicesdo not take into consideration how cells will be migrating into thescaffolds as well as how their presence influences the surrounding,uninvolved tissue. Additionally, prior art scaffolds did not take intoconsideration the joint fluid and how it impacts maturation andmaintenance of healthy hyaline cartilage. Although prior art syntheticscaffolds, whether single phase, multi-phase, or of gradientconstruction have proven suitable for growth and maturation of cellswithin a bioreactor, these prior art devices are unsuitable for directimplantation, for at least the reasons that follow.

Applicants have made the surprising discovery that in effecting therepair of cartilage defects, prior art synthetic implants and syntheticbi-phasic implant devices failed to recognized the need to ignore thenormal histological and mechanical gradient of the articular cartilage,and instead focused on the limited cell population surrounding thedefect and its slow rate of tissue formation within the devicesresulting from this sparse population of cells. The prior art syntheticimplants mistakenly focused on speeding up the rate of cell migrationwithin the scaffold in hopes of getting tissue to form rapidlythroughout the device prior to collapse of the scaffold. This increasedrate of cell migration was done using chemotactic ground substances suchas hyaluronic acid, cell seeding or biologics. All this served to do wasto spread out the cell population and reduce the rate of hyalinecartilage tissue formation, and as a result, biased any new tissuegrowth of cartilage towards the fibrocartilage lineage. Although somesuccess in establishing hyaline cartilage can be seen in small defectsof 5 millimeters or less, larger defects show tell tale signs ofcollapse or dimpling in the center of a repair plug, as the lessdesirable fibrocartilage, which has grown within the prior art devices,succumbs to the forces within the joint. Additionally, prior art devicesshow a halo or ring of collapsed tissue around the periphery of thedevice do to lack of intimate contact with the uninvolved tissue thathas retracted away from the defect site.

Another discovery of applicants is that prior art devices do not addressthe instantaneous articular cartilage tissue contraction that occurswhen the surface of hyaline cartilage is cut or torn. Upon damage, thecartilage retracts way from the defect site forming a funnel. Thus priorart devices, upon implantation, do not make contact with the surroundinguninvolved cartilage.

The uninvolved host tissue, that is, the normal tissue adjacent to andsurrounding the defect site that is not involved with the defect, isable to influence the activities of cells that migrate into andestablish themselves at the periphery of a scaffold placed into thedefect. The cells of the uninvolved tissue, along with the extracellularmatrix of the uninvolved host tissue adjacent to the periphery of theimplanted scaffold are already established as hyaline cartilage and thusmechanically and chemically react to stresses appropriately. Through aprocess, sometimes referred to as mechanical signal transduction, theestablished host tissue is able to influence the phenotype andextracellular matrix produced by the adjacent cells in the scaffold thusproducing the desired hyaline cartilage. Specifically, cartilaginoustissues perform specialized functions under normal physiologicalconditions. Anomalous mechanical loading of these tissues often leads topathology. For example, the lack of mechanical stimulation of a jointleads to suppression of proteoglycan synthesis and release of mediatorsresponsible for degradation of cartilage matrix components. This isbelieved to be the cause of collapse or dimpling of the newly formedcartilage seen with prior art devices.

The molecular mechanisms controlling the response of cartilaginoustissues to their mechanical environment are not completely understood.Furthermore, there is a dearth of knowledge about the modes ofmechanical signal transduction in chondrocytes. Several theoriesconcerning the molecular mechanisms through which mechanical stimulimodulate the expression of cartilage extracellular matrix (ECM)components have been proposed, some of which are: 1) receptor mediatedcell-ECM adhesion contributes to the transduction of mechanical signalsin chondrocytes, 2) mechanical signal transduction in chondrocytesrequires activation of the phosphoinositol and/or cyclic AMP (also knownas Cyclic adenosine monophosphate or cAMP) signaling pathways, and 3)mechanical stimulation of the expression of aggrecan is mediated throughactivation of specific cis-acting elements of the promoter and/or UTRs(untranslated regions) of the aggrecan gene. No matter the specificmechanism through which it happens, applicants believe that theinfluence uninvolved host tissue has over the cells in the scaffoldmatrix extends approximately 2.5 millimeters. Thus, this places a limitof success for prior art devices having a matrix equal to, or less stiffthan the surrounding host tissue to 5 millimeters in diameter. However,any device having a cartilage scaffold matrix greater in stiffness thanthe surrounding host tissue will not be properly influenced bymechanical signal transduction and will either form calcified tissues ordisorganized fibrocartilage that collapses as the matrix degrades andthe tissue experiences stress loading.

In order to prevent the observed central collapse or dimpling within thecartilage layer of prior art implants, applicants have discovered that anew type of scaffold must be made that retards rapid migration of cellsacross the entire diameter of the device, thereby concentrating cellsand cell activity at the edges of the device, promoting rapid andsystematic tissue conduction and maturation, moving from the outer edgeof the device towards the interior. Additionally, the area within thecartilage region of the scaffold where cell activity is occurring mustbe less rigid than the surrounding uninvolved tissue, to ensure that itis subject to the mechanical influences of the adjacent uninvolvedtissue.

Within the bone layer, known prior art devices failed to recognize theimpact a rigid scaffold has on the surrounding uninvolved tissue.Whereas malleable elastic scaffolds (scaffolds that can be deformed andthen return to their original shape) are desirable for the cartilagelayer, rigid stable scaffolds (scaffolds that resist deformation) arerequired for proper migration and attachment of bone forming cells.However, nearly the opposite conditions are required for stability ofexisting bone, as micromotion is beneficial to healthy bone structure.Micro-motion and/or stresses are necessary to keep healthy bone frombecoming osteopenic. Osteopenia refers to bone mineral density that islower than normal. Bone mineral density has been shown to drop inhealthy individuals who are bedridden, as well as in astronauts who havereduced stress on their skeletal system due to the effects of reducedgravity while in space. As this occurs, the bones lose minerals,heaviness (mass), and structure, making them weaker and increasing theirrisk of collapse and or breaking. Localized bone mineral density losshas been witnessed due to stress shielding caused by orthopedic rods andplates. During repair of damaged cartilage with prior art devices, voidsand osteopenic zones have been observed to form below implanted tissuescaffolds. The theory behind this pathology formation is that stressshielding, caused by the presence of porous tissue scaffolds, results inbone density loss. The scaffolds dampen vibrations that would normallybe transferred through the malleable elastic articular cartilage to thecalcified region and then conducted deeper into the bone. Theseconductive forces are necessary for normal bone biology. The conductedforces in normal bone located below an articulating joint travel notonly through the bone trabecula, but also through the viscous gel ofbone marrow surrounding the bone trabecula. This is because the bonetrabecula located under the cartilage of a joint shows a generalhistologic pattern of elongated channels radiating out from thecalcified region into the subchondral bone. Thus forces are not onlytransmitted down the rigid walls of the channels formed by thetrabecula, but are also transmitted by the gelatinous bone marrowcontained within the channels. Two functional problems identified withrigid porous scaffolds of prior art devices are as follows. First theserigid devices do not contain elongated channels and thus they tend todissipate and dampen the hydrostatic pressure pulses that would normallyflow through viscous fluids. Secondly these devices are too rigidthrough the cartilage region thus not allowing for initial compressionto establish a pressure wave through the bone marrow.

In order to prevent undesirable bone voids from forming in uninvolvedtissues adjacent to the repair device, what is needed is a scaffoldcapable of transferring forces through the device, and into the tissue.This deep bone mechanical stimulation is due to compression of thearticular cartilage region generating mechanical and fluidic forcesduring normal movement in the joint.

Concerning the synovial fluid, prior art devices fail to recognize therole this substance plays in maintaining healthy articular cartilage.Synovial fluid is a thick, stringy fluid found in the cavities ofsynovial joints. Synovial fluid reduces friction between the articularcartilage surfaces as well as providing cushioning during movement. Theinner membrane of synovial joints is called the synovial membrane and itsecretes synovial fluid into the joint cavity. This fluid forms a thinlayer (about 50 microns thick) at the surface of cartilage and seepsinto the micro-cavities and irregularities in the articular cartilagesurface, filling all empty space thus presenting a uniform, smoothsurface. The fluid in the articular cartilage effectively serves as asynovial fluid reserve, during movement; the synovial fluid held in thecartilage is squeezed out mechanically to maintain a layer of fluid onthe cartilage surface. This so called weeping lubrication ensures thatincreased friction does not occur as some of the lubrication fluid isswept away during joint movement.

Synovial tissue is composed of vascularized connective tissue that lacksa basement membrane. Two cell types (type A and type B) are present:Type B cells produce synovial fluid. Synovial fluid is made ofhyaluronic acid and lubricin, proteinases, and collagenases. Synovialfluid exhibits non-Newtonian flow characteristics. The viscositycoefficient is not a constant, the fluid is not linearly viscous, andits viscosity increases as the shear rate decreases.

Almost all of the protein constituents of synovial fluid are derivedfrom plasma. The passage of plasma proteins to synovial fluid is relatedto the size and shape of the protein molecule. Most proteins withmolecular weights less than 100,000 daltons are readily transferred fromone fluid space to another. Thus synovial fluid is a plasma dialysatemodified by constituents secreted by the joint tissues. The majordifference between synovial fluid and other body fluids derived fromplasma is the high content of hyaluronic acid (mucin) in synovial fluid.Normal synovial fluid contains 3-4 mg/ml hyaluronan (hyaluronic acid), apolymer of nonsulfated polysaccharides composed of D-glucuronic acid andD-N-acetylglucosamine joined by alternating beta-1,4 and beta-1,3glycosidic bonds. Hyaluronan is synthesized by the synovial membrane andsecreted into the joint cavity to increase the viscosity and elasticityof articular cartilage and lubricates the surfaces between synovium andcartilage. Both fibroblasts beneath the synovial membrane intima andsynovial membrane-lining cells produce this mucopolysaccharideconstituent of synovial fluid.

Synovial fluid is believed to have two main functions: to aid in thenutrition of articular cartilage by acting as a transport medium fornutritional substances, such as glucose, and to aid in the mechanicalfunction of joints by lubricating the articulating surfaces. Articularcartilage has no blood, nerve, or lymphatic supply. Glucose forarticular cartilage chondrocyte energy is transported from theperiarticular vasculature to the cartilage by the synovial fluid.Synovial fluid contains lubricin secreted by synovial cells. Synovialfluid is chiefly responsible for so-called boundary-layer lubrication,which reduces friction between opposing surfaces of cartilage. There isalso some evidence that synovial fluid helps regulate synovial cellgrowth. Synovial fluid serves many functions including: reducingfriction by lubricating the joint; absorbing shocks; and supplyingoxygen and nutrients to, as well as removing carbon dioxide andmetabolic wastes from, the chondrocytes within articular cartilage.

Normal synovial fluid does not clot but may exhibit thixotropy, theproperty of certain gels to become fluid when exposed to shear forcessuch as shaking. On standing at room temperature, normal synovial fluidmay assume gelatin-like appearance, characterized by higher viscosities.When shaken it will assume a normal fluid nature. Many enzymes have beenfound in the normal synovial fluid. Alkaline phosphatase, acidphosphatase, lactic dehydrogenase, and other enzymes are present indetectable quantities. Enzymes enter the synovial fluid directly fromthe plasma or may be produced locally by the synovial membrane orreleased by synovial fluid macrophages. Synovial fluid also containsphagocytic cells that remove microbes and the debris that results fromnormal wear and tear in the joint.

Some prior art devices utilize fluid impermeable layers at the cartilagesurface, the bone/cartilage interface, or both locations, or have rigidarticular cartilage regions resistant to receiving fluid from thesynovial space. These types of structures serve as barriers that preventthe normal transfer of essential elements from the synovial fluid, intoand out of the cartilage region. What is needed is a device capable offacilitating joint fluid therapy to the chondrocytes within the defect.Joint fluid therapy encompasses delivering, receiving, accumulating andcontrolling the location of desirable factors or molecules present inthe synovial fluid while also delaying or preventing destructivefactors, such as digestive enzymes, from prematurely degrading thematrix. These desirable factors or molecules can be those naturallyoccurring within the synovial fluid or biologically active agentsadministered into the synovial fluid.

SUMMARY OF THE INVENTION

This invention includes implantable biphasic devices for the repair oftissues of a living being, especially, cartilaginous tissue defects. Inthe embodiment of a biphasic device, the device has a first region and asecond region, each being specific for the growth of a particular tissuetype. In an embodiment useful for repair of cartilage defects, the firstregion is specific for cartilage tissue growth, and the second region isspecific for bone growth.

In one aspect of the invention, the device is an electro-kineticimplant, in which at least a portion of the device features twojuxtaposed materials that form a malleable matrix, where the firstmaterial presents a positively charged surface, and the second materialpresents a negatively charged surface. As the malleable matrix isdeformed under the application of pressures, such as may occur whileimplanted in a living being, an electrical potential is produced as aresult of interactions, and interruptions, between the charged surfacesof the first and second materials. In one embodiment, the malleablematerial will be malleable while hydrated, though it may be rigid, or atleast capable of being handled without deformation, while in a drystate. In another embodiment, the malleable material may exhibit anelastic property, tending to return to its original shape after havingbeen deformed. The first material of the malleable materials may be aparticulate, especially a fibrous particulate, and the second materialof the malleable material may be a hydrogel, such that the particulateis suspended within the hydrogel, and upon deformation, the hydrogel andparticulate move relative to each other. The malleable material may beporous. The materials may be ceramics, natural polymers, syntheticpolymers, or combinations thereof.

The charges in the charged surfaces may be the result of exposure of theconstituent materials to acidic or basic environments, plasma gas, or aresult of the attachment of charged substances to the materials.

In one embodiment, the first and second materials of the malleablematerial are collagen, with the first collagen material, such as afibrous collagen, presenting a positively charged surface, and thesecond collagen material, such as a hydrogel presenting a negativelycharged surface. In this embodiment, the charged surfaces of the fibrouscollagen and hydrogel collagen may be created by exposing each of thecollagens to solutions, where one collagen is exposed to a solutionhaving a pH above the isoelectric point of the collagens, and the othercollagen is exposed to a solution having a pH below the isolectric pointof the other collagen.

Another aspect of the invention provides for the transmission of forcesand loads throughout a malleable matrix component making up at least aportion of the implantable device. In one embodiment, the malleablematrix component is created having a first and second material, wherethe first material is a hydrogel and the second material is aninterconnected network of fibers. In this embodiment, the hydrogelcomponent may be collagen, or hyaluronic acid, and the fibrous componentmay be collagen or chitosan. The malleable matrix component is able toprovide joint fluid therapy to the cells or tissue within the implantdevice as it is arranged to transmit forces throughout the entire, or atleast substantially the entire volume, of the malleable matrixcomponent, as forces applied will cause a vortex ring or gyre due to theinteractions of the interconnected fibers pulling on each other, as theyare displaced within the hydrogel material. It is believed that thethree-dimensional transmission of forces throughout the malleablematerial will result in the malleable material, or at least the hydrogelcomponent of the malleable material, receiving and accumulatingdesirable factors or molecules from surrounding fluids, which may beutilized by cells within the device.

In one embodiment, the malleable material is one phase of a biphasicdevice, and corresponds to the cartilage region, thus the malleablematerial may be attached to a rigid base corresponding to the boneregion.

In another aspect of the invention, the implant provides for thesystematic tissue conduction and growth from the surrounding cartilagetissue, and retards the formation of tissue in the interior of theimplant. In this manner, it is believed that the growth of the incorrecttype of tissue can be avoided, and better ensure that only the desirablehyaline cartilage is formed. In an embodiment, the device may comprise agradient, where the gradient is arranged to retard the tissue formationmost at or near the center of the implant (when viewed top down), andtransitions to little or no retardation of tissue formation towards theperimeter of the implant, adjacent to normal cartilage tissue. Thegradient may be in the form of a circular gradient, and may be uniformthroughout the device from upper surface to lower surface, oralternatively may vary from top to bottom. The gradient may be a smoothtransition or gradual gradient, or alternatively a stepwise gradienthaving well defined regions within the gradient. The gradient may be aconcentration gradient, such as biologically active agents, additives,or combinations thereof. The gradient may be a physical gradient, suchas porosity, density, expansion, swelling, elasticity, hardness,compressibility, and combinations thereof. The gradient may be amaterial gradient, or chemical gradient, such as molecular weight,cross-linking, hydrophobicity, polarity, crystallinity, and combinationsthereof. The gradient may be part of the first phase of a multiphasicdevice, and corresponds to the cartilage region, and may be attached toa rigid base corresponding to the bone region.

In another aspect of the invention, the multiphasic implant provides forthe transmission or conduction of pressure forces through the device,down to the underlying bone tissue below the device; in this manner,bone tissue loss below the device, such as may occur due tostress-shielding, may be minimized or avoided. One embodiment of animplant device capable of transmitting such forces would present a boneregion presenting a porous material and a rigid penetrating forceconductive material capable of transmitting the forces received from amalleable cartilage region to the underlying tissue. The forces to betransmitted may be hydrostatic and directed through channels runningthrough the bone region material, or alternatively force transmissionmay be in the form of kinetic pressure pulses through the rigidconductive material arranged in the bone phase. The rigid conductivematerial may be in the form of columns arranged perpendicular to the topand bottom surfaces of the implant, and may flare out to a widerdimension at the junction with underlying bone. The rigid conductivematerial may be in the form of a rigid multi-facetted web structureoriented perpendicular to the top and bottom surfaces of the implant. Inanother embodiment, the rigid conductive material is a wedge or conethat transmits the forces through the implant to the underlying bone,but may also transmit some forces laterally as an outward force to theporous bone region material.

In yet another embodiment the multiphasic device capable of transmittingpressure forces presents at least a first material in the form of atleast two porous rigid scaffolds, where the first material is separatedby at least a second material in the form of a malleable elastichydrogel, and where the hydrogel is capable of transferring hydrostaticpressure pulses through the bone region of the device in order toprevent bone voids from forming in external underlying bone tissue.

The various embodiments described herein may be at least partially orcompletely resorbed by the living being. The various embodimentsdescribed herein may also feature drugs, biologically active agents, orother additives in all or at least a portion of the device.

Various medical uses of the above-described invention are describedbelow. Other features or advantages of the present invention will beapparent from the following drawings and detailed description of theinvention, as well as from the claims.

DESCRIPTION OF THE DRAWINGS

FIG. 1 is a perspective depiction of a circular gradient.

FIG. 2 is a cross-section depiction of the circular gradient of FIG. 1.

FIG. 3 is a perspective depiction of a circular gradient having atapered construction from upper surface to lower surface.

FIG. 4 is a perspective depiction of a circular gradient having anhour-glass shape, wherein the gradient zones are wider in the upper andlower surfaces, and featuring a narrow mid-section.

FIGS. 5 and 6 are perspective depictions of multiple circular gradientswithin the same device,

FIG. 7 is a cross-sectional depiction of a biphasic device as found inthe prior art, having a cartilage region comprising a gel or porousmaterial.

FIG. 8 is a cross-sectional depiction of a biphasic device, having acartilage region arranged as a web or matrix, where the web or matrix isable to telegraph applied forces through substantially all of thecartilage region, by the movement of the web or matrix constituents in amanner analogous to a vortex ring, or gyres.

FIG. 9 is a 1-year histology slide of a repair site having had a priorart biphasic implant device implanted, after the device has beencompletely absorbed, wherein stress shielding is evident.

FIG. 10 is a cross-sectional depiction of an implant embodiment that isarranged to transmit forces or loads through the device to underlyingtissue below, using a rigid central column.

FIGS. 11 a and b are cross-sectional depictions of another implantembodiment arranged to transmit forces or loads through the device usinga stiff, multifaceted web structure and filler porous material.

FIG. 12 is a cross-sectional depiction of an implant embodiment that isarranged to transmit forces or loads through the device using a rigidcentral conically shaped center post.

FIG. 13 is a depiction of a multi-layered cylinder comprising variousmaterial thicknesses and densities, where the layers can serve totransmit forces or loads through the device to underlying tissue below.

FIGS. 14 a and b are depictions of a multi-layered cylinder materialhaving swellable properties upon hydration, and capable of transmittingforces hydraulically through the device to the underlying tissue below.

DETAILED DESCRIPTION OF THE PREFERRED EMBODIMENTS

A device and methods are disclosed for treating tissue deficiencies,defects, voids and conformational discontinuities produced by congenitaldeformities, tissue pathology, traumatic injuries and surgicalprocedures, particularly those located in mammalian bone and cartilage.In one embodiment, the device is to provide the means by which hyalinecartilage tissue can be conducted across a tissue specific firstscaffold region by controlled migration of chondrocytes and/or cartilageprecursor cells. Additionally, in an embodiment, the scaffold region canbe designed to affect the concentration, location and activity offluids, factors, molecules or other biologically active agents receivedfrom, or delivered to, the extracellular fluids, especially synovialfluid. Thus, the device provides means to regenerate a first specificform of tissue.

A tissue specific second scaffold region may be attached to the firstregion for controlled migration of osteoblasts and/or bone precursorcells. Thus described, an embodiment of the device is a biphasic device,wherein the device consists of two main parts, the cartilage region, andthe subchondral bone region, which are joined at an interface surface.Additionally, an embodiment provides a means for deep bone mechanicalstimulus by conduction of mechanical and/or fluid forces originating in,or being applied to the cartilage specific scaffold region. Thesestimuli will be conducted through the subchondral bone region into theadjacent uninvolved subchondral bone.

In a bi-phasic embodiment, the cartilage region can be joined or boundto the subchondral bone region of the device by a number of processes,including but not limited to, heat fusion, heat welding, adhesives,glues or solvent welding. The resulting union between the twoarchitectural regions is preferably very strong and can withstand anyhandling required to package the device as well as any forces deliveredto it as a result of the implantation technique without permanentlydistorting the device's internal architecture of void spaces.

The interface surface between the two regions may be a permanent ortemporary barrier to the passage of cells, fluids, or biologicalcomponents (e.g. growth factors, proteins, cells signals, etc.) so longas it does not interfere with the transmission of mechanical stimuliresulting from compression of the first region.

In the biphasic device embodiment, the ingrowth or formation of tissuewould be specific to the device region, that is, cartilage tissue wouldgrow into the cartilage region of the device, and bone tissue would growinto the bone region of the device based upon the cells within theimmediately adjacent tissues, as well as mechanical and chemical signalsprovided by the individual layers of the device. Furthermore, each ofthe cartilage regions and bone regions may provide for a physicalstructure that is appropriate to the type of tissue for which it isproviding a substrate. That is, the bone region will provide a stablesubstratum for attachment of bone or bone forming cells, while thecartilage region will provide a malleable elastic substratum capable ofallowing the surrounding uninvolved tissue to mediate, or affect, thecompression and motion of the scaffold adjacent to the host tissue.Additionally, additives capable of enhancing the growth of the targettissue are contemplated within the current invention. Additives in thebone region can include ceramics, glass, glass-ceramics, bioactiveglasses, as well as biologically active agents. Additives in thecartilage region can include gelatinous materials, as well asbiologically active agents. The additives may be initially loaded intothe cartilage region for interaction internally within the device and/orfor external device delivery. Additionally, additives can originatewithin the synovial fluid and be passively or actively transported intothe cartilage region of the device. Non-limiting examples of materialsand additives useful in construction of the various embodiments of thedevices described herein can be found in Table 2.

The architecture of each device region may be formed utilizingestablished techniques widely practiced by those skilled in the art ofmedical grade polymers. These methods may include injection molding,extrusion and machining, vacuum foaming, precipitation, sintering,spinning hollow filaments, solvent evaporation, soluble particulateleaching or combinations thereof. For some methods, plasticizers may berequired to reduce the glass transition temperature to low enough levelsso that polymer flow will occur without decomposition. Additionallyadditives such as plasticizers or particulates can be added to polymersto make them more or less malleable (malleable materials can be elasticas defined earlier or plastic wherein they do not return to thereoriginal shape after deformation) in order to provide the desiredmechanical properties for the specific device region they will belocated in. For example, a normally rigid polymer may incorporate aplasticizer to make it malleable and thus useful in the cartilage regionwhereas rigid particles could be added to a malleable polymer to providea stable substratum suitable for use in the bone region.

In an embodiment, the osteochondral repair device will be formed as aplug, typically circular in cross-section, and shaped to fill a void ordefect created through the cartilage layer and into the underlying bone.Additionally, it is recognized that the plug may have a tapered form,such that one end of the device is larger than the other. A defectsuitable for accepting the device can be created in a manner known tothose skilled in the art, for example, using the device as described inU.S. application Ser. No. 11/049,410, or alternatively using defectcreation techniques known as the OATS procedure. It is recognized thatalternative shapes other than cylinders, may be utilized, for examplejoining or overlapping circular elements together into one larger shapewill allow for larger defect areas to be repaired with coring tooldevices suitable for smaller defects (e.g., approximating an oval,figure eight or a cloverleaf shape). Additionally, non-circular shapesmay be utilized as well, such as by providing plug devices withalternative cross-sections, for example, polygonal shapes may beemployed or combined (e.g. rectilinear, triangular, hexagonal, etc.), asthe polygons may be joined alongside other devices to form a mosaiccovering a larger area than could a single device.

Once there is a void created in the bone to accept the implant device(e.g., core created by a coring tool), the implant device is preparedfor implantation. The implantable device may be directed into the voidthrough arthroscopic means, or alternatively by hand into the exposedbone void. Preferably, the device is loaded into an insertion tool.Though any known insertion tool or mechanism may be employed, it isenvisioned that the delivery can be accomplished with an insertion toolincluding a device-containing barrel with a delivery end, and also aplunger extending into the barrel for ejecting the device out thedelivery end, in a manner similar to a wide mouth syringe. The insertiontool is then placed adjacent to the opening, or directed into theopening, and the device is then ejected from the delivery tool, into thebone void. Preferably, care is taken, both in the creation of the void,and in the delivery of the device, to avoid damaging the healthy nearbytissue, particularly the cartilage tissue and chondrocytes.

Once cellular tissue is fully established within the defect repair site,it is expected that normal loads will be fully supported by the newtissue. For biodegradable devices, the device degrades and is eventuallyresorbed or removed from the implantation site. This occurs as thedevice is degraded and provides for the complete transfer of loadbearing ability from the device to the ingrown tissue, prior to thedevice's load bearing ability falling below the levels required to aidin tissue incorporation. Within this document, biodegradable,degradable, bioresorbable, resorbable, bioerodable and erodeble may beused interchangeably.

The various embodiments of a tissue repair device as described hereinmay be implanted dry, or hydrated with biologically relevant fluids, forexample, saline, blood, bone marrow aspirate or Platelet Rich Plasma(PRP). Also, growth factors, hormones, drugs, cells or other usefulbiologically active agents, can be used to hydrate the device. Thesematerials can provide therapy to the cells migrating into the implant,the surrounding tissue, or the synovial fluid. Optionally, growthfactors, hormones, drugs, cells or other useful biologically activeagents can be located within the synovial fluid and adsorbed into theimplant by passive or active means. For reference, a non-exhaustive listof biologically active agents that may be incorporated into at least aportion or the entirety of the various embodiments contained herein canbe found in Table 1.

In healthy osteochondral tissues, for example a knee, having a verticalaxis that is in the load bearing direction, and a horizontal axis thatis normal to both the tissue surface and the load bearing direction,typically, the encountered loads due to natural movement and gravity areable to be transmitted or conducted through the soft tissues of thejoint, and into the hard bony tissues. The load transmission is largelyvertical, being in the direction of load application, and createscompression of the soft tissue, however, due to the interconnectivity ofthe soft tissues, particularly across the transverse layer of thearticular cartilage, some portion of the loads are distributed laterallyas well, to adjoining soft tissue. One effect of this lateraldistribution is that a force of a given magnitude, having been appliedat only a small area at the top of the soft tissue, and beingtransmitted through the soft tissue, would result in the force beingdistributed over a wider area at the bottom of the soft tissue, and intothe bone. Given the wider distribution of the force over a larger area,a compressive force in only a small area of the articulating surface canprovide deep bone mechanical stimulus to a large area of subchondralbone, with the peak force felt directly below the originatingcompressive force and lesser amounts of conductive stimulus radiatingoutward.

In a similar fashion, where there has been a defect in osteochondraltissue, and a plug device is implanted, the loads that would havenormally been transmitted by healthy tissue, would now desirously betransmitted by the plug device as well. Consequently, not only should adevice that is inserted into a defect beneficially be able to withstandthe expected loads in the defect location, both in the direction of theinitial force application, and also laterally as the force isdistributed through the soft tissue, but should also be able toadequately transmit or conduct those forces through the device and intohealthy adjacent tissues.

Where the device is bioresorbable and also supports the growth of newtissue, it is beneficial to ensure that the degradation characteristicsof the device are such that new tissue ingrowth is structurallycompetent, meaning that it is able to support the expected loads in thedefect area, at least coincidentally, or prior, to the degradation ofthat portion of the device being subsumed by the new tissue ingrowth. Inthis manner, the device can avoid the dimpling failure mode seen inprior art devices, as a portion of the device becomes structurallyincompetent, the newly grown and structurally competent tissues canprovide the required weight bearing ability as well as the ability totransmit mechanical stimulus.

One embodiment is intended to address the previously described dimplingfailure modes, where, it is believed, a portion of the repaired defectarea collapses prior to the growth of structurally competent tissue. Itis believed that the collapse manifested as dimpling at the surface ofthe repair site, is a result of failure in either, or both of, theremaining structures of the implanted device, or in the new tissueingrowth replacing the device as it degrades. This embodiment alleviatesthis occurrence by providing for a resorbable implant structure thatfosters satisfactory and controlled tissue ingrowth, and provides forthe last invaded and absorbed portion of the device to be degraded afterthe tissue growth in the device is able to withstand and transmit theencountered loads, also termed “structurally competent”. This maygenerally be achieved in one of two broad manners. One may ensure thatthe device has adequate structural competence for a period of time thatis long enough to allow adequate tissue restoration prior to the devicebecoming structurally incompetent. Alternatively, one may accelerate theradial ingrowth of competent tissue into the device, such that cells aresignificantly established and forming the morphologically correcttissue, thereby creating structurally competent tissue in a shorter timeframe, prior to the device losing its structural competence. There arevarious techniques that may be employed for achieving each of thesegoals, such as controlling porosity, density, cross-linking, drugdelivery, cell seeding, etc. These techniques will be discussed later.It is recognized that one or more techniques may be combined into asingle device, to create an ideal solution.

With reference to the following figures, applicants will describevarious embodiments for presenting a tissue repair device. FIGS. 1-6depict various gradient formations that may be employed within acartilage region of a biphasic device, that could allow for competenttissue growth to be achieved as the device is degraded and ultimatelyabsorbed, thereby avoiding a mechanical failure of the device caused bycollapse or dimpling of the central portion of the newly establishedtissue within the cartilage region.

In one embodiment, and with reference to FIG. 1, a device is providedhaving at least one controlled gradient in the device that is arrangedconcentrically around a vertical axis and normal to the cartilagesurface. This circular gradient may provide, for example, for a longerduration of implant structural competence as tissue grows inconcentrically (or in the form of accelerated tissue regeneration fromthe outer zone) and spreads to the inner zone depicted at the center ofthe device and whereas the zones depicted in the outer portion of thedevice allows for rapid cell invasion and the inner central zones retardcell invasion, or extracellular matrix deposition, until such a time asthe cells in the outer zones have laid down the appropriateextracellular matrix, influenced by the mechanical reaction to loadingof the uninvolved adjacent cartilage, in the form of hyaline cartilage.It is important to prevent the occurrence of tissue formation asisolated islands, which are not in contact with the uninvolved normalarticular cartilage, as isolated islands will not receive appropriatemechanical stimulus from the surrounding uninvolved tissue.

The controlled circular gradient in the device of FIG. 1 is termed a“bull's-eye” gradient. The “bull's-eye” gradient refers to the way thedevice appears when viewed from the cross-sectional direction depictedin FIG. 2. As can be seen, the bull's-eye gradient consists of a centralcore region or zone 300, surrounded by one or more annular rings. Inthis depiction, there are two annular rings, 100 and 200, concentricallyarranged about the central core. Though it is recognized that more orless annular rings could easily be achieved as well. The controlledgradient depicted in FIG. 1 with the cross section as shown in FIG. 2 isuniform throughout the length of the device. It is also contemplatedthat the gradient could vary along the vertical axis, for example,differing in dimension to provide non-uniform cross-sections throughoutthe length of the device, as can be seen, for example in FIG. 3 and FIG.4. In FIG. 3, the gradient has two annular rings 110, and 210,surrounding a core region 310. In FIG. 4, the gradient has two annularrings 120, and 220, surrounding a core region 320. It is recognized thatthe gradients depicted by the figures may exist within a separatestructural element in the form of a cylinder or disk

Gradients can fall into many different groups including but not limitedto concentration, chemical, physical and material. The invention can beprovided in a great variety of useful shaped devices, as will bediscussed later, where the gradients of the invention may be created byvarying one or more of a variety of characteristics, including porosity,density, molecular weight, cross-linking, hydrophobicity,hydrophilicity, polarity, drug concentration, drug delivery, material,expansion, swelling, elasticity, hardness, compressability,crystallinity, cell seeding, etc. To provide further clarity, selectcharacteristics will be explored more fully below, with reference toFIG. 1, as the simplest embodiment, however, it is recognized that othershapes or gradient forms for practicing the invention could employsimilar characteristic or compositional gradients.

Controlling the density of specific regions of the device may be usefulto provide greater structural resistance to compressive loads. In anembodiment, a gradient can be constructed where the center of the devicehas a higher density then the outer edge. The density change may beachieved, for example, by varying any of the porosity, pore size or porenumber in each region of the device, or by varying the molecular weightof the polymer in various zones. For the example of a bull's-eyegradient, as depicted in FIG. 1, the device may provide higher densitypolymers or less porous scaffolding at the center zone 300, and thenfurther removed from the center to the perimeter on the cross-sectionalplain of the device, the material becomes less dense and more porous.This embodiment with high porosity at the outer zone 100, allows for thecells to migrate quicker initially at the outer zone 100, but retardstheir ability to reach the central zone 200 and inner zone 300, thusconcentrating the cells in the outer zone 100. Central zone 200 willhave a porosity or density in between that of the interior and exteriorof the device. This will also extend the duration of structuralcompetence at the core, as the mechanical strength of the core iselevated due to the increased density, and can thus be tailored to stayabove a minimum threshold value (as determined by the expected physicalloading in the defect area) for a longer period of time, as the devicegoes through biological degradation. The increased duration ofstructural competence at the center zone 300 allows more time for tissueto infiltrate, grow, and become structurally competent in the core ofthe device, prior to the total degradation or structural collapse. Thoseskilled in the art will recognize other types of gradients that can beused to decelerate cellular migration, as will be discussed.

An embodiment of the device may provide for a gradient by usingbiologically active agents (e.g., drugs, cells, growth factors, etc),ceramics, glass, metals or polymers, all of which are included in theterm “additives” incorporated into the device. In this embodiment, theouter zone 100 of the device may provide an elevated additiveconcentration, relative to the additive concentration provided at thecentral zone 300 of the device. For the specific case of growth factors,or other agents that will enhance cellular chemotaxis and growth, thishigh concentration in the outer zone 100 will help recruit cells to theouter edge of the device faster and can increase tissue regeneration atthe exterior of the device, resulting in a shorter time period to reachstructural competence as the new tissue continues to grow into themiddle zone 200, and then into the core zone 300.

Controlling the rate of cross-linking of the polymer in specific regionsof the device may be useful to provide greater structural resistance tocompressive loads. In an embodiment, a gradient can be constructed wherethe innermost zone 300 of the device has a higher percentage ofcross-linked polymers than the outer zone 100 with middle zone 200having a percentage of cross-linked polymer somewhere in-between. As aresult of the cross-linking, the polymer will be more stable underloads, and less subject to biodegradation and bioresorption, resultingin a longer duration of structural competence in the more extensivelycross-linked regions, relative to the lesser cross-linked regions of thedevice. This increased resistance to compressive loads will protect anycells prematurely gaining access to the core portion of the cartilageregion from receiving incorrect mechanical signals prior to beinginfluenced by the encroaching tissue. Cells receiving little to nomechanical stimulus will either attempt to move down the bone lineageline (i.e., differentiate), or if isolated from high oxygen content asnaturally occurs in the articular cartilage, will remain relativelydormant while waiting for mechanical or chemical stimulus. In this waythe innermost, more cross-linked region will not inadvertently allowcells to commit to the bone or fibrocartilage line, but instead causethe cells to wait to be influenced by the mechanical properties of thetissue being conducted through the matrix from the outer zones as thematrix degrades and becomes softer. With reference to FIG. 1, the corezone 300 may be a highly cross-linked polymer, and transition to outerzone 100 that is not cross-linked at all, or features lesscross-linking. As stated previously, mechanical signal transduction iscritical to differentiation of the newly forming tissue, any devicehaving a cartilage scaffold matrix greater in stiffness than thesurrounding host tissue will not be influenced by mechanical signaltransduction and will either form calcified tissues or disorganizedfibrocartilage that collapses as the matrix degrades and the tissue isstress loaded. Thus it is important to initially concentrate the tissueforming cells in the outer zones where they can be influenced by thesurrounding uninvolved tissue while at the same time preventingpremature collapse of the central zone. The device as described hereinis intended to set up the best circumstances to allow for the formationof the correct tissue type.

Controlling the compositional makeup of specific regions of the devicemay be useful to provide regions with longer durations of structuralcompetence. In an embodiment, a gradient can be constructed bycontrolling the polymer blend ratio in each of the zones to providevarying mechanical strength, or degradation rates. For example, theinnermost zone of the device may be manufactured with a polymer or ablend of polymers that provides enhanced resistance to degradation, orincreased mechanical strength, when compared to the polymer, or blend ofpolymers provided in the outer zone of the device. In this embodiment,the center core of the embodiment will feature an enhanced duration ofstructural competence relative to the outer zone of the device

As a specific non-limiting example, and with reference again to FIG. 1,natural polymers such as collagen may be used to create regions withvarying durations of structural competence. The outer zone 100 of thedevice can be constructed from soluble collagen that posses no fibersand is gelatinous by nature. This allows for more rapid cellular tissuein-growth to the outer region of the device as the collagen has a lowcompressive modulus and/or degrades at a rapid rate allowing the newlyrecruited cells to be stimulated by the mechanical forces necessary tolay down the appropriate tissue matrix. The middle region 200 of thedevice may be constructed from fibrillar collagen. Being of a higherhierarchical structure the fibrillar collagen provides greaterstructural integrity and/or greater resistance to degradation. In thecore zone 300 the collagen may be fibrous, thereby providing evengreater mechanical properties and/or greater resistance to degradationthan either of the outer zones. Thus, using the hierarchical structureof collagen, a gradient can run through the spectrum of gelatin, solublecollagen, fibrillar collagen, fibrous collagen and collagen in the formof decellularized tissue, with or without its extracellular matrixcomponents, some or all of which can be cross-linked as a tool forfurther control. Additionally, the gradient could be based on lengthand/or thickness and/or density of fibrils or fibers. For instance ahomogenous soluble collagen disk may contain an additive such ascollagen fibers with the mass or density of said collagen fibersincreasing as one proceeds or travels from outer zone 100 towards innercore 300. Collagen gradients, as well as other material gradients, mayalso be the result of differing animal sources (bovine, porcine, equine,etc), or use of genetically engineered collagen, for instance from plantsources.

Regions with varying durations of structural competence may also beachieved with different types or species of polymers from natural orsynthetic sources. As an example, outer zone 100 can be made fromhyaluronic acid, which is very easily degradable, while the middleregion 200 can be constructed from natural polymer that is moreresistant to degradation such as collagen. The inner core 300 maycontain an even tougher polymer such as chitosan. Non-limiting examplesof materials and additives useful in construction of devices describedherein can be found in Table 2.

It is recognized that various embodiments of the device may provide morethan one gradient, examples of which are depicted in FIGS. 5 and 6. Inthese multi-gradient embodiments, a pair of gradients are created, afirst bull's-eye gradient may extend from its widest dimension at theupper surface, and as one travels down the vertical axis, the bull's-eyeof the first gradient is shown to diminish in cross-section, ultimatelycontracting to a point where the zones merge. The second gradient mayextend from the lower surface, and diminish in area as one travels upthe vertical axis.

Specifically with regard to the multiple gradient embodiment, asdepicted in FIG. 5, the first gradient has two annular rings 130, and230, surrounding a core region 330, and the second gradient has twoannular rings 140 and 240, surrounding a core region 340. As can beseen, the second bull's-eye gradient has its largest dimensional area atthe lower surface, and as one travels up the vertical axis, thebull's-eye gradient forms around the cone formed by the first bull's-eyegradient described previously. Thus the outer dimension of the secondgradient is shown to remain uniform, while the inner zone of the secondgradient forms as an annular ring surrounding the cone of the firstgradient. As one nears the upper surface of the device, the secondbull's-eye gradient regions merge into a narrow annulus, preferably ator near the upper surface of the device.

Specifically with regard to the multiple gradient embodiment, asdepicted in FIG. 6, the first gradient has two annular rings 150, and250, surrounding a core region 350, and the second gradient has twoannular rings 160 and 260, surrounding a core region 360. As can beseen, the second gradient is created as an inverse to the firstgradient, and has its largest dimension at the lower surface, and as onetravels up the vertical axis, the bull's-eye of the second gradient isshown to reduce in cross-section, ultimately reducing to a point wherethe zones merge. In this embodiment, the first and second gradients arecomposed of unrelated characteristics or materials, and the presence ofone gradient will not necessarily interfere with the presence of theother, thus they can be seen to overlap or extend into each other asdepicted in FIG. 6. It should be recognized that these gradients mightexist within a device having the gross shape in the form of a cylinder.For example, a plug device having a uniform porosity with one of thegradients being a first biologically active agent, and the othergradient being a second biologically active agent. In another example,one of the gradients may comprise a structural gradient (e.g., density,cross-linking, etc.)

A potential application of this “reverse-cone” of FIG. 5, or“inverse-cone” of FIG. 6 is that the gradients can be employed tooptimize the balance required between promoting rapid cell regenerationand tissue competence, against the required need for adequate mechanicalcompetence of the device as well as regulating the rate of devicedegradation, which is so important to the success of the device. In thisembodiment, it is believed that the first gradient (upper) couldpreferably be a density gradient, such as can be created by controllingthe porosity, pore size, pore density, or polymer molecular weight, andthe second gradient (lower) could preferably be an additive gradient(e.g., growth factors, drugs, ceramics, cells, etc.)

Yet another “bull's-eye” design can have a narrow mid-section creatingan hourglass look, as depicted in FIG. 4. (It should be noted that thisdepiction only represents the gradient, such as mechanical integrityresulting from fiber incorporation, and that the general matrix in whichthe gradient resides is not pictured.) The cross-section in this area ofthe device gradient is much smaller, relative to the upper and the lowerregions of the device gradient, but still maintains the “bull's-eye”pattern. In such an embodiment, the mechanical integrity of the deviceis maintained by the gradient depicted. Thus the gradient may not alwaysoccupy the entirety of the device to be implanted. This extra area notpictured can be filled, for example, with highly porous polymers thataren't required to provide any structural competence properties, andwhose main objective would be to receive host cells and thus promotemore rapid tissue regeneration in the external regions closest to theuninvolved host tissue. That is, for the external regions where littleor no structural competence is required to be provided by the deviceallowing the uninvolved adjacent host tissue to mechanically influencethe region, it is preferable to provide a material that maximizes theamount and extent of cellular ingrowth into the exterior of the device,in order to provide a foothold of structurally competent tissue withinthe device as quickly as possible.

For controlled gradients generally, it is contemplated that the gradientbe formed by altering some material or property within the device in amanner corresponding to the patterns depicted in the figures. Startingfrom the innermost zone at the core and transitioning through theintermediate zones out to the outer region, the gradient would providesome characteristic that varies as one moves further out from thecenter. For the sake of simplicity and ease of visualization, much ofthe explanation in this application only discusses the example of FIG.1, however, it is recognized that the teachings of this application alsoare applicable to the other examples and figures contained in thisapplication as well.

As depicted in FIG. 1, the gradient may feature zones delineated by theconcentric annular rings that provides a recognizable or detectableborder or interface between each of the differing zones presented byFIG. 1. Alternatively, it is recognized that a continuous transitionalgradient or gradual circular gradient could be provided, where there isa gradual change in the characteristic, from the core region andprogressing out to the outside circumference, and the rings depicted inFIG. 1 are merely representative of the direction of the transition.

It is envisioned that a device providing for the various gradientcharacteristics described herein could be manufactured as an intactdevice, using carefully controlled lyophilization techniques forcreating these gradients. Alternatively, a series of components may bemanufactured, each varying in a particular characteristic. Subsequently,the components may be shaped to a form, where each component will formone of the zones, and thereafter be assembled into a final device. Forexample, and with reference to FIG. 1, a core piece could bemanufactured, and later inserted into annular rings sizedconcentrically, where each of the assembled components will create thegradient desired in the final device. Alternatively, one component maybe provided as a scaffold material in the manufacture of the othercomponents, thereby forming a multi-zoned device providing a gradientcharacteristic. An example of this manufacturing method would includethe injection molding of a central skeleton followed by theincorporation of other less dense open-celled matrices whose densitiesprogress from the central structure outward towards the perimeter of thefinished device.

It is also envisioned that gradients could be made or created bycompressing a starting porous polymer matrix to collapse or sacrificepores and thus develop a device having the various zones as previouslydescribed. In addition, these gradients could be developed by startingwith granulated material, and then through the use of heat andcompression, could yield a finished device containing varying porositiesand physical shapes. For example, fine granular material having anaverage diameter less than 50 microns can be placed in the center of acylindrical mold creating a central core. Around this can be pored amedium granular material having an average diameter in the range of50-100 microns creation a middle zone. A course granular material havingan average diameter exceeding 100 microns in turn will surround this.Compression and heat may then be used to fuse this granular materialtogether to create a bull's-eye gradient device.

It is also contemplated that that the cartilage region of the currentinvention could be made to expand after implantation. In this manner,the device would provide intimate contact with the surroundinguninvolved cartilage tissue that has retracted away from the defecthole, as the removal of a circular defect from normal articularcartilage has been observed to result in differential retraction of theedges. Depending on the depth of the defect, the edges retract more inthe superficial zone as compared to the deeper zones after a circulardefect is removed with a punch. Normal human cartilage, with an intactsuperficial zone, curls when removed from the underlying bone. Theretraction away from the defect site, as well as the curling of theremoved cartilage, is the result of the high tension existing within thesuperficial zone of articular cartilage. This results in a cone orfunnel shape forming in the articular cartilage portion of a surgicallycreated defect, narrowing as one moves down towards the subchondral boneportion of a surgically created defect. The current inventionanticipates this and thus can be capable of radial expansion in order toensure a tight fit. For example, a cylindrical device can be place intoa newly created defect and expand until is has a shape as shown in FIG.3.

Applicants have made an additional surprising discovery that ineffecting the repair of cartilage defects, prior art synthetic implantsand synthetic bi-phasic implant devices failed to recognize theimportance of synovial fluid in the maintenance and repair of articularcartilage. As an additional consideration in the development of a devicefor repair of articular cartilage one needs to understand how friction,cyclic motion, electric potential and synovial fluid all work togetherto maintain the articular cartilage phenotype. Under normalphysiological conditions, articular cartilage provides a nearlyfrictionless surface between moving joint. To help lubricate thesejoints, the body uses synovial fluid. This fluid component consistsprimarily of water with dissolved solutes and mobile ions.

Solute transport in biological tissues is a fundamental process of life,providing nutrients to cells and carrying away waste products. Inavascular tissues such as adult articular cartilage, solute transportoccurs primarily across the articular surface, with synovial fluidmediating exchanges with the synovium lining the joint capsule. Aprimary mechanism of solute transport is through diffusion. Themechanism of passive diffusion in healthy cartilage has been shownexperimentally to be enhanced by cyclical loading of the cartilage, andby electro-osmotic flow both, of which mechanisms lead to convectiveflow within the tissue. Other avascular tissue types that respondsimilarly to articular cartilage include tendon, ligament, meniscus andannulus thus the techniques described herein for use in cartilage repairby manipulating the natural fluid and electric potential in the regionmay be used on these other tissue types as well. It is also envisionedthat these techniques could be beneficial on vascularized tissue thatare elastic in nature, including but not limited to blood vessels andskin.

Within cartilage, it is recognized that the synovial fluid acts as atransport medium for substances into and out of the articular cartilageregion. This is necessary because the articular cartilage region is anon-vascular tissue. Substances are transported into and out of thearticular cartilage region due to repetitive mechanical stimulusfollowed by a period of rest. During active mechanical stimuli,molecules located within the synovial fluid are actively transportedinto the articular cartilage layer. This allows the concentration ofmolecules within the cartilage tissue to exceed that of the synovialfluid. During rest, the concentration will return to equilibrium. Inthis way, necessary substances located within the synovial fluid areforced into the cartilage tissue, whereupon the cells can absorb them.Waste products are excreted by the cells into the interstitial space ofthe tissue where they build up. During a period of rest, the systemmoves towards equilibrium and thus the waste products move out of thecartilage tissue and into the synovial fluid wherein they are ultimatelytransported into the vasculature and away from the knee.

Thus, vital nutrients are supplied to the non-vascular or avasculartissues from the blood vessels located at the margins of the tissue. Thetransport of nutrients through the dense complex extracellular matrix tothe cells making up these tissues relies mainly on diffusion. Poornutrient supply has been suggested as a potential mechanism fordegenerative processes that affect the avascular tissues(i.e.—osteoporosis, disk degeneration, etc.) and is also suspected infailure of prior art cartilage implants.

The effects of dynamic compression on chondrocyte biosynthesis have beenwell characterized in cartilage explants and chondrocyte-seededscaffolds. In explants, continuously applied dynamic compression anddynamic tissue shear have been found to increase synthesis of proteinsand proteoglycans.

Studies of articular cartilage metabolism have demonstrated that staticloading, as well as loading below a characteristic frequency of 0.001Hz, leads to biosynthetic inhibition, whereas dynamic loading stimulatestissue synthesis. This enhanced biosynthetic response is believed toresult from an enhanced nutritional supply, as well as a tissuebiosynthetic response under dynamic loading, and thus resulting inenhanced fluid flow and changes in cell shape or mechanotransduction.Static compression of articular cartilage has been shown to reduce thediffusivity of various solutes within the tissue, and has beenimplicated in the altered biosynthetic response of the tissue to staticloading. Growth factors, which have been shown to regulate thebiosynthetic response of articular cartilage, are generally largesolutes with molecular weights on the order of tens of kilodaltons. Afurther benefit of dynamic loading is growth-factor uptake. It has beenshown that dynamic compression accelerates the biosynthetic response ofcartilage to free IGF-I and increases the rate of transport of freeIGF-I into the cartilage matrix, suggesting that cyclic compression mayimprove the access of soluble growth factors.

Dynamic compression, thus, augments the transport of solutes inavascular tissues such as cartilage. However, the effect of mechanicalcompression on the distribution and metabolism of nutrients is difficultto directly evaluate. To this end, research has been conducted onsynthetic gels in order to answer these questions.

Exposing an agarose gel, submerged in a fluid medium containing targetmolecules, to repetitive mechanical compression can crudely simulate thedynamic tissue compression system. It has been observed that althoughthe target molecules move against the concentration gradient onto thegel, they are not evenly distributed throughout the gel. The moleculesonly move into the area under direct mechanical stimulus. If it was thecase that cartilage tissue behaved identically, then it would followthat cells around the edges of the cartilage, would be deprived ofnecessary substances. However, as will be discussed, cartilage does notbehave identically to agarose gel, though it does exhibit the similarphenomenon of increasing the concentration of molecules as a result ofrepeated compression. This unequal distribution of necessary substancesis a shortcoming of prior art devices having a gel-like property withinthe cartilage region. Normal articular cartilage overcomes this unequaldistribution by having a dense fibrous layer, known as the transverselayer that absorbs and distributes mechanical stimulus across theentirety of the tissue layer by providing a mechanical coupling of thecartilage molecules to each other. In this way, necessary substances areactively moved into the entirety of the cartilage tissue layer.

Similar to the normal cartilage tissue layer, a preferred form of thecurrent invention allows for uniform incorporation of necessary targetmolecules by providing a biodegradable, insoluble malleable elastic gelor hydrogel like substratum containing a sufficient concentration offibers so that they form a penetrating interconnected phase. The gel orhydrogel can also present an interconnecting porosity. The fibers,making up a second phase can be entangled, entwined, interwoven, knittedor in some other fashion connected in a three-dimensional web or matrixso that stresses in the form of a push or pull are telegraphedthroughout the entire device. In this way the current invention iscapable of receiving joint fluid therapy throughout its entire volume.

FIG. 7 represents prior art biphasic device 700 having a cartilageregion of simple pores, or in the form of a gel. When force 710 isapplied to the surface of cartilage region 720, pressure waves 730remain focused just below original force 710. Contrasted with theembodiment of FIG. 8, which depicts a cross-section of biphasic device800 wherein connected fibers 810 are located within layer 820. Whenforce 830 is applied to the surface of layer 820, downward pressureforce 830 causes connected fibers 810 to pull on each other,telegraphing pressure force 830 throughout the entire volume of layer820, creating a circular force, vortex, vortex ring, toroid, or gyre, asrepresented by arrows 840. It should be noted that although shown in asingle plane, the described circular force occurs three-dimensionallyestablishing a vortex ring, that is, multiple vortexes or gyres withinthe device. The potential vorticity of fluid within the device isdirectly related to the volume of displacement within the device matrixfrom the downward pressure force. In the simplest sense, vorticity isthe tendency for elements of the fluid to “spin.” and can be related tothe amount of “circulation” or “rotation” in the fluid contained in thematrix caused by the gyres. As new host tissue grows into the edges ofthis embodiment of the device, forces applied to cartilage tissuedistant from the device will be transmitted through the host tissue andinto the device.

The cartilage layer of an embodiment of the device will be composed ofat least two phases. This first phase is an insoluble gel or hydrogelcapable of adsorbing and concentrating target molecules from thesynovial fluid when placed under repetitive compressive forces. Thesecond phase will be a fibrous component associated with or containedwithin the gel phase having sufficient connectivity so that acompressive force applied to one location of the cartilage layer istransmitted throughout substantially the entire volume of the cartilagelayer. In order to achieve this the minimum average fiber length forfibers randomly located within the gel should be approximately equal tothe thickness of articular cartilage, which is from 2-3 millimeters inlength. The maximum average fiber length should not exceed 1.5 times thediameter of the devices so as to prevent curling or coiling of thefibers preventing them from being taut within the matrix and thusdampening the transmission of mechanical stimulus. These same lengthrestrictions apply to interwoven or knitted type devices in as much asconnecting nodes or knots holding the structure together should be nocloser than 2-3 millimeters apart and no farther apart than 1.5 timesthe diameter of the devices. For the example of a plug implant devicehaving a diameter of 10 mm, and a cartilage region thickness of 3 mm,the average length of the fibers would be in the range of 2-15 mm, andthe average distance between connecting nodes or knots would be in therange of 2-15 mm.

The material phases can be fabricated from natural and/or syntheticpolymers including but not limited to collagen, elastin, keratin,chitosan, hyaluronic acid, silk, alginate, polyethylene glycol (PEG) andcombinations thereof. (Non-limiting examples of materials and additivesuseful in construction of devices described herein can be found in Table2.) One or more of the phases can also contain biologically activeagents such as those listed in Table 1.

The biologic activities of the chondrocyte population are regulated bygenetic, and other biologic and biochemical factors, as well asenvironmental factors. It has often been noted that physicalenvironmental factors, such as stress, fluid flow, electric fields, etc.are as strong as biologic factors in regulating cellular activities.There has been much research on the effects of mechanical and/orhydrostatic/osmotic pressure loading on cartilage explant metabolism.Such studies have been specifically aimed at elucidating possible“mechano-signal” transduction (also referred to as“mechanotransduction”) mechanism(s) that might govern the chondrocytes'biosynthetic activities in maintaining and organizing the extracellularmatrix (ECM) comprising the tissue. Over decades many researchers haveobserved electrical events in cartilage, but few studies have focused onthe details of the electrical potential within the ECM where thechondrocytes reside. This phenomenon of electromechanical orelectrokinetic cell signaling has also be ignored by prior art devices.Electromechanical or electrokinetic cell signaling is not to be confusedwith mechanotransduction, as mechanotransduction does not createelectrical potential.

The electromechanical signals that chondrocytes perceive in situ are theresult of stresses, strains, pressures and the electric fields generatedinside the extracellular matrix when the tissue is deformed. Thepotential induced by convection in the presence of a pressure gradientin deformed tissue is the “streaming potential”. The potential inducedby diffusion in the presence of a concentration gradient in statictissue is the “diffusion potential”.

Avascular tissues such as cartilage are composed of water, collagenenmeshed in a proteoglycan gel, and various matrix proteins. The osmoticpressure of these tissues is mainly due to the high density of chargedcarboxyl and sulfate groups on the glycosaminoglycans of theproteoglycans within the tissues. When avascular tissues are deformedunder loading, interstitial fluid flow occurs, even though the hydraulicpermeability of the tissues is very low. The electrical response of thetissues also changes when it is compressed due to the effects ofdiffusion potential and streaming potential.

The diffusion potential is the electric potential caused by theseparation between the bulk positive and bulk negative charges caused bythe gradients of mobile ions within the different fluid regions of thetissue or between the tissue fluid and the synovial fluid.

Streaming potential is defined as the difference in electric potentialbetween a diaphragm, capillary, or porous solid and a liquid that isforced to flow through it. The definition of streaming potential canalso include the difference in electric potential caused by theoscillation, separation or flow of a gel in relationship to a diaphragm,capillary or porous solid. Specifically, it is the potential that isproduced when a liquid or gel is forced to flow through a capillary or aporous solid. The streaming potential is one of four relatedelectrokinetic phenomena that depend upon the presence of an electricaldouble layer at a solid-liquid/gel interface. This electrical doublelayer is made up of ions of one charge type that are fixed to thesurface of the solid and an equal number of mobile ions of the oppositecharge which are distributed through the neighboring region of theliquid/gel phase. In such a system the movement of liquid/gel inrelation to the surface of the solid produces an electric current,because the motion of the liquid/gel causes a displacement of the mobilecounterions with respect to the fixed charges on the solid surface. Theapplied potential necessary to reduce the net flow of electricity tozero is the streaming potential. Streaming potential is related to zetapotential by factors that include the electrical conductivity andfluid/gel viscosity. The value of streaming potential under givenconditions of conductivity and pressure can be used to judge howstrongly the tissue will interact with anionic or cationic molecules.The zeta potential is a good predictor of the magnitude of electricalrepulsive force. A resulting voltage is measured between electrodeprobes on either side of a boundary. This voltage is then compared withthe voltage at zero applied pressure.

The source of electrical events, as measured on the outside surface ofnormal articular cartilage, derives from the fixed, immobilized ortrapped negative charges ˜SO3 and COO2, distributed along thechondroitin, keratin sulfates and hyaluronan molecules comprising theaggrecan inside the extracellular matrix of the tissue. Theseproteoglycans may be assumed to be “immobilized and trapped” inside theextracellular matrix, and therefore considered to be fixed to theextracellular matrix. Together with the surrounding collagen network,these proteoglycan macromolecules form the cohesive, strong,porous-permeable, charged, collagen/proteoglycan solid matrix. By virtueof the electro-neutrality law, there is always a cloud of counter-ions(e.g., Ca, Na) and co-ions (e.g., Cl) dissolved in the interstitialwater surrounding the fixed charges in the extracellular matrix. Theseions may move by convection with the interstitial fluid due to ahydraulic pressure or by diffusion through the fluid due to aconcentration gradient or by conduction, drifting through the fluid as acurrent due to an electric field. Forces for the electric current insidethe tissues include the mechano-chemical force generated by the gradientfrom movement of ions resulting from compression and diffusion caused byion concentration gradients.

Within deformable tissues such as articular cartilage, intervertebaldisk, epiphyseal (growth) plate, and cornea, the electric fieldsresulting from mechano-chemical forces are constantly present. Thus, forsuch tissues, both streaming potential and diffusion potential mustalways exist inside the tissue and in fact they always compete againsteach other. The streaming potential arises from the slight separation ofthe bulk of the positive charges from that of the negative charges dueto the flow convection effects caused by a pressure gradient fromdeformation of the tissue. The diffusion potential arises from theslight separation of the bulk of positive charges from that of thenegative charges due to diffusion caused by the gradients of mobileions. It is believed that electrical events inside the tissue areimportant in stimulating chondrocyte biosyntheses. It is also believedthat non-uniform electrical effects resulting from deformation occurswhen a tissue is softened during a disease process such asosteoarthritis. In osteoarthritic cartilage, with matrix degradation,the intrinsic compressive stiffness always diminishes, thus affectingchondrocyte deformation and metabolic activities as well as the natureof the mechano-electrochemical events within cartilage when it isdeformed.

Another preferred embodiment of the current invention presents acartilage region that takes into consideration both diffusion potentialand streaming potential in its constructions. The cartilage layer ofthis preferred device will be composed of at least two phases. Thisfirst phase is an insoluble gel or hydrogel capable of adsorbing andconcentrating target charged molecules from the synovial fluid whenplaced under repetitive compressive forces. The second phase will be afibrous component contained within the gel phase having sufficientconnectivity so that a compressive force applied to one location of thecartilage layer is transmitted throughout the entire volume of thecartilage layer. This allows creation of a disparity between the overallcharges of the synovial fluid from that of the cartilage layerestablishing the diffusion potential. In addition to this it isdesirable for the first phase to predominantly contain either positiveor negative charges while the second phase will predominantly containcharges opposite that of the first phase. In this way a pressuregradient from deformation of the cartilage layer of the preferredembodiment creates a slight separation between the charges of the firstphase from that of the second phase, as the gel and fibers flex, thusestablishing the streaming potential. If desirable, one or both phasescan be cross-linked. Thus the electric potentials created by such anembodiment simulate that which occurs in normal articular cartilage,thus improving and/or stimulating chondrocyte biosyntheses and thusarticular cartilage tissue formation.

In one possible method for the manufacture of an embodiment that takesinto consideration both diffusion potential and streaming potential,insoluble collagen fibers are exposed to a more basic chemicalenvironment (above the pH of the collagen's isoelectric point) in orderto bring the surface of the collagen above its isoelectric point andthus providing a predominantly negative charge to the surface of thefibers composing the second phase of the devices. These negativelycharged fibers are embedded within a collagen gel or hydrogel that wasexposed to a more acidic chemical environment (below the pH of thecollagen's isoelectric point) so as to drive this collagen below itsisoelectric point to provide a predominantly positive charge to thisfirst phase. This is unlike prior art devices that contain two phases ofcollagen wherein both collagens are on the same side of the isoelectricpoint.

In another embodiment, biodegradable polyester fibers (ie -PLA, PGA,PCL, etc), which have been subjected to surface modifications, such asexposure to acids, bases, or plasma gas processes) are used in thesecond phase of the device.

In another embodiment, hyaluronic acid gel or hydrogel having apredominantly negative charge is used as the first phase thatencapsulates and surrounds a second phase of chitosan fibers having anoverall positive charge. When making combinations such as hyaluronicacid and chitosan, care must be taken so that a polyelectrolytic complex(PEC) is not formed as this will not allow the charges to separateduring compression and thus no electric potential will occur.

In another embodiment, an electrically neutral hydrogel first phaseenvelops a charged fibrous second phase, wherein the gel allows mobileions to penetrate and take up residence within the gel thus balancingout the charge of the fiberous second phase. As described previously,deformation of the combined matrix will result in charge separation,creating the electric potential. An example of an electrically neutralhydrogel would be a PEC. Such a PEC could be manufactured by varioustechniques known in the art, incorporating known components. The neutralhydrogel PEC could be created by the combination of charged components,such as hyaluronic acid—chitosan, collagen—chitosan, and hyaluronicacid—collagen.

It is also recognized that the second phase material can be composed ofparticulate materials that are not fibrous or polymeric in nature solong as they provide the necessary charged surface. A non-limiting listof materials suitable for this use can be found in table 2.

Those skilled in that art will identify other combinations of positivelyand negatively charged materials all of which are embraced by thisdisclosure for use in creation of an electro-kinetic tissue repairdevice.

As already discussed, in some embodiments, part of the function of thedevice is to transfer forces or loads, experienced by the cartilagelayer, through the devices and into the subchondral bone. This deep bonemechanical stimulus is necessary to prevent stress shielding thatcurrently results in bone voids below the device. FIG. 9 shows 12-monthhistology from a prior art device that provided stress shielding to theunderlying bone. Box 910 shows the approximate location of the implantthat has completely resorbed. Soft tissue void 920 within the bone isthe result of this stress shielding.

Both cartilage and bone are living tissues that respond and adapt to theloads they experience. If a joint surface remains unloaded forappreciable periods of time the cartilage tends to soften and weaken.Further, as with most materials that experience structural loads,particularly cyclic structural loads, both bone and cartilage begin toshow signs of failure at loads that are below their ultimate strength.Research into bone healing has shown that some mechanical stimulationcan enhance the healing response and it is likely that the optimumregime for a cartilage/bone graft or construct will involve differentlevels of loading over time in order to properly repair a damagedregion. This same observation was concluded by Surgeon Julius Wolff backin the 19^(th) century and is still known today as Wolff's law.

Many prior art implants that are made for use in repairing damaged boneand cartilage are fabricated from soft materials and deform when theyare implanted into a cored hole in the bone. These implants do notprovide a means for the transfer of loading through the implant forstimulating the growth of new bone at the bottom or side walls of thecored hole, or even controlling or preventing osteopenia orosteoporosis. Other implants that are fabricated as bone void fillersare made from rather stiff materials such as ceramics. These devices canprovide a means for mechanical stimulation; however, the implant must beprecision fitted to the bone void in order to create the proper lengthto match up with the hole that has been cored into the patient's bone.Since any protrusion of these devices will result in higher contactpressure, which may further damage the cartilage in joint areas, it isnot advisable to use these devices for cartilage repair.

For osteochondral transplantation involving the replacement of damagedcartilage sites with harvested plugs taken from the patients' joint, itis also difficult to match the cored hole depth with the exact implantlength. This is a function of the design of the coring tool as well asthe technique utilized by the surgeon. For some coring tools, the coredhole will exhibit a very uniform cylindrical shape, however, the bottomsurface may be inconsistent and have a rather jagged and irregularsurface. This can create gaps or void pockets under the implant orcreate a void between the top of the implant and the mating rotatingbone and prevent any transfer of forces or pressure during the healingprocess. In addition, the surgeon is concerned about protrusion of theharvested plug creating too much pressure on the transplanted hyalinecartilage thereby damaging this tissue as the joint moves. Therefore,the surgeon often creates a deeper recipient site defect then the lengthof the harvested plug. This allows the surgeon to control the finalposition or height of the implanted device; however, this is assumingthat the frictional forces alone will provide enough stability for theplug to stay in position. This also creates a void space under theimplant, which prevents contact from occurring with the subchondralbone.

Other studies have shown that bottoming out the implant can provide forbetter support and stability during the time that the cells are growinginto the newly implanted plug. However, bottoming out the implant cancause high compressive forces during insertion, which can also damagethe transplanted cartilage during the surgery. These same studies havealso shown that these implanted plugs are more stable and can be cut toshorter lengths if they are bottomed out.

In order to obtain loading through the cartilage/bone region of anydevice, contact and pressure are required to exist. As previouslydiscussed, it may not be possible to create a tight enough fit betweenthe implanted device and the cored hole in the patient's bone.Therefore, the implanted device needs to provide the capabilities toexpand and contract to fill this space.

Based on these requirements, it is envisioned that a device could bedesigned so that a portion of it has the ability to expand and contractlike an extension spring. Once the device is implanted into a coredhole, the expansion and contraction of the implant would provide theproper functionality. In addition, it is desirable to also createsufficient contact with the walls of the cored hole.

A cartilage/bone repair device is envisioned which takes intoconsideration the transfer of structural loads or pressures that may beseen by the implant once it is installed into a cored-out hole in therecipient's bone.

In various embodiments, the implant may be made of different materialsor different forms of the same material. As an example, a rigid supportskeleton can be injection molded from a PLA polymer and this samepolymer can be chemically processed to create an open-celled foamstructure. Both of these materials would act in completely differentways in regards to their absorption characteristics, their load transfercharacteristics, and their biological cell attraction characteristics.

In other embodiments, the implant may include various means of securingitself within the area of bone repair. These securing means can includemechanical methods such as teeth or ridges that are incorporated aroundthe outside surfaces of the device. These teeth or ridges can alsoassist with the transfer of forces through the device and into thesurrounding bone.

In further embodiments, the device could utilize differentcharacteristics formulated into the structural make up of the device inorder to promote the take up of fluid thereby causing a hydraulic effectin a portion of the device, which would create a means of expansion andthereby allow for pressure to be transferred through the device.

In another embodiment, the device contains fluid swellable expansionzones that provide for a tight fit within the void and allow formicro-motion while other porous stable zones allow for cell attachmentand tissue growth.

Various methods can be utilized for transferring the forces or loadsthrough the device in order to provide mechanical stimulation to thebone interfacing surfaces. As shown in FIG. 10, device 1000 has centercolumn 1010 positioned under cartilage layer 1020 that transcends downthe center and then transitions to a larger diameter at the bottom toallow the transfer of force or pressure between the upper surface of theimplant and the implant/bone interface layer at the bottom of thedevice. Porous matrix 1030 surrounds center column 1010 and makescontact with the host tissue. Center column 1010 can be porous, but isrigid and thus conductive of mechanical stimulus that would be dampenedby porous matrix 1030. It is preferable that porous matrix 1030 swellsshortly after placement into the tissue void so direct contact is madewith the tissue void walls. In addition transitioning to a largerdiameter at the bottom, center column 1010 can also transition to alarger diameter at the top (not shown), presenting an hourglass type ofshape. Additionally, center column 1010 can be formed from a smalldiameter cylinder with a thin flat plate on the bottom and optionallythe top (not shown). The porosity, if present, in center column 1010 canbe random, or in the form of elongated channels capable of conductinghydraulic forces.

FIG. 11 a shows device 1100 having multi-facetted web structure 1110that is oriented perpendicular to the top and bottom surfaces of device1100. In this configuration, the web acts as a stiffener to transfer theload originating in cartilage layer 1120 through the implant. Secondarymaterial 1130 is a less dense, more porous structure formed in betweenthe spokes of web structure 1110. FIG. 11 b shows a top down view withthe cartilage layer removed so that the relationship of the spokes ofweb structure 1110 and secondary material 1130 can easily be visualized.In this embodiment, the web would continue to transfer the forces intothe subchondral bone region while bone growth was occurring within theporous structure of secondary material 1110 found in between the webs orspokes of web structure 1110. As bone growth completed the encroachmentof this area, it would assist with the load or pressure transfer whilethe materials of web structure 1110 started its degradation and eventualremoval. Web structure 1110 can have holes or slots within its structureto allow intercommunication of the secondary material 1130.

As shown in FIG. 12, device 1200 has conically shaped center post 1210sitting below cartilage layer 1220. Center post 1210 wedges into outercylinder layer 1230 possessing a shaped inner cavity designed to receivecenter post 1210. Center post 1210 may extend completely through outerlayer cylinder layer 1230 as pictured or may instead just come flush tothe base of device 1200. The tapered shape of center post 1210 providesfor a means of seating the implant while also providing a method fortransferring mechanical stimulus to all sections of the subchondral boneregion. When downward force 1250 is applied to device 1200 outercylinder layer 1230 is experiences outward force 1240 thus providingimproved seating of device 1200 into a cored bone void. Thus forcesapplied to cartilage layer 1220 pass into center post 1210 and aretransferred to the tissue void.

FIG. 13 is composed of a multi-layered cylinder containing variousmaterial thicknesses and densities. The layers can be constructed to actto transfer the pressure between the top surface of the device and thebottom surface. The composition of these various layers can also beutilized to create hydraulic swelling to thereby create the spring-likeeffect previously described.

FIG. 14 a shows a simplified example of a bone region multi-layeredcylinder 1400. To simplify understanding, the cartilage layer is notpictured. Swellable layers 1420 separate rigid porous layers 1410. Moreor less layers are also contemplated. Upon implantation or exposure toliquid, swellable layers 1420 imbibe fluid and become an uncompressible,flexible hydrogel as depicted in FIG. 14 b where rigid porous layers1410 are now separated by swollen layers 1430. Referring to FIG. 14 b, aforce applied to the cartilage layer (not shown) is transferred as apressure wave through the device so long as rigid porous layers 1410 donot exceed 4 mm in thickness and have an average porosity greater than50 microns and are rigid enough to avoid collapse of their porosity thusnot dissipating the pressure wave prior to it reaching the bottom layerand finally being conducted into the underlying bone. Optionally one ormore holes, 2 millimeters in diameter or greater can exist in layers1410 allowing pillars of hydrogel to connect swollen layers 1430. It maybe I that newly forming bone needs a stable substratum to attach to sothat bone forming cells can lay down extracellular matrix. Bone formingcells known as osteoblasts are approximately 50 microns in diameter andshould establish themselves in newly forming islands of boneapproximately 1 mm in diameter, thus the minimum thickness of porouslayer 1410 is 1 millimeter. The thickness of swollen layer 1430 has nomaximum, but should be at a minimum of 5 microns with a preferredthickness of 50 microns to trap a sufficient amount of fluid and thusfunction as an incompressible hydrogel capable of transferring pressurewaves.

In another embodiment (not shown) porous particles having a minimumapproximate diameter of 1 millimeter can be surrounded by a swellablematerial wherein the swellable material maintains connectivitythroughout the entire device. In this way, pressure waves and micromotion, necessary for establishing bone external to the device, can beconducted through the swellable material matrix while the porousparticles provide a stable platform for attachment and proliferation ofosteoblasts. As a non limiting example, porous particles composed ofceramic, polymer or composites of the two can be suspended within ahydrogel forming material such as collagen, hyaluronic acid, chitosan,alginate, keratin, or PEG. In addition to being a homogenous material,the hydrogel can be formed into a porous network presenting fluidswollen struts or partitions defining fluid containing pores.

The bone region of all the above devices can be designed so that theyprovide the required expansion and transfer of force as the materialsdegrade. This transfer of force can occur through the use of rigidpolymeric or ceramic elements, incompressible hydrogels or combinationsthereof. As more cells are stimulated to grow into the implanted matrix,newly formed tissue will help to continue the transfer of the mechanicalstimulation.

The inclusion of groups and subgroups in the tables is exemplary and forconvenience only. The grouping does not indicate a preferred use orlimitation on use of any material therein. For example, in Table 1, thegroupings are for reference only and not meant to be limiting in anyway. Additionally, the tables are not exhaustive, as many other drugsand drug groups are contemplated for use in the current embodiments.There are naturally occurring and synthesized forms of many therapies,both existing and under development, and the table is meant to includeboth forms.

Numerous other embodiments and modifications will be apparent to thoseskilled in the art and it will be appreciated that the above descriptionof a preferred embodiment is illustrative only. It is not intended tolimit the scope of the embodiments contained herein, which are definedby the following claims. Without further elaboration the foregoing willso fully illustrate our invention that others may, by applying currentor future knowledge, adopt the same for use under various conditions ofservice.

TABLE 1 Examples of Biologically Active Agents Adenovirus with orwithout genetic material Angiogenic agents Angiotensin Converting EnzymeInhibitors (ACE inhibitors) Angiotensin II antagonists Anti-angiogenicagents Antiarrhythmics Anti-bacterial agents Antibiotics ErythromycinPenicillin Anti-coagulants Heparin Anti-growth factors Anti-inflammatoryagents Dexamethasone Aspirin Hydrocortisone Antimicrobial AntioxidantsAnti-platelet agents Forskolin Anti-proliferation agents Anti-rejectionagents Rapamycin Anti-restenosis agents Antisense Anti-thrombogenicagents Argatroban Hirudin GP IIb/IIIa inhibitors Anti-virus drugsArteriogenesis agents acidic fibroblast growth factor (aFGF) angiogeninangiotropin basic fibroblast growth factor (bFGF) Bone morphogenicproteins (BMP) epidermal growth factor (EGF) fibringranulocyte-macrophage colony stimulating factor (GM-CSF) hepatocytegrowth factor (HGF) HIF-1 insulin growth factor-1 (IGF-1) interleukin-8(IL-8) MAC-1 nicotinamide platelet-derived endothelial cell growthfactor (PD-ECGF) platelet-derived growth factor (PDGF) transforminggrowth factors alpha & beta (TGF-.alpha., TGF-beta.) tumor necrosisfactor alpha (TNF-.alpha.) vascular endothelial growth factor (VEGF)vascular permeability factor (VPF) Bacteria Beta blocker Blood clottingfactor Bone morphogenic proteins (BMP) Calcium channel blockersCarcinogens Cells Adipose cells Bone marrow cells Blood cellsEndothelial Cells Epithelial cells Skeletal muscle cells Smooth musclecells Stem Cells Umbilical cord cells Fat cells Bone cells Mesenchymalstem cells Progenitor cells Cartilage cells Cellular Material Bonemarrow Cells with altered receptors or binding sites FibroblastsGenetically altered cells Glycoproteins Growth factors Lipids LiposomesMacrophages Reticulocytes Vesicles Chemotherapeutic agents (e.g.Ceramide, Taxol, Cisplatin) Cholesterol reducers Chondroitin CollagenInhibitors Colony stimulating factors Coumadin Cytokines prostaglandinsDentin Etretinate Genetic material Glucosamine Glycosaminoglycans GPIIb/IIIa inhibitors L-703,081 Granulocyte-macrophage colony stimulatingfactor (GM-CSF) Growth factor antagonists or inhibitors Growth factorsAcidic fibroblast growth factor (aFGF) Autologous Growth Factors Basicfibroblast growth factor (bFGF) Bone morphogenic proteins (BMPs) BovineDerived Growth Factors Cartilage Derived Growth Factors (CDF)Endothelial Cell Growth Factor (ECGF) Epidermal growth factor (EGF)Fibroblast Growth Factors (FGF) Hepatocyte growth factor (HGF)Insulin-like Growth Factors (e.g. IGF-I) Nerve growth factor (NGF)Platelet Derived endothelial cell growth factor (PD-ECGF) PlateletDerived Growth Factor (PDGF) Recombinant NGF (rhNGF) Recombinant GrowthFactors Tissue Derived Cytokines Tissue necrosis factor (TNF)Transforming growth factors alpha (TGF-alpha) Transforming growthfactors beta (TGF-beta) Tumor necrosis factor alpha (TNF-.alpha.)Vascular Endothelial Growth Factor (VEGF) Vascular permeability factor(UPF) Growth hormones Heparin sulfate proteoglycan HMC-CoA reductaseinhibitors (statins) Hormones Erythropoietin Immoxidal Immunosuppressantagents inflammatory mediator Insulin Interleukins Interlukin-8 (IL-8)Interlukins Lipid lowering agents Lipo-proteins Liposomes LipidsLow-molecular weight heparin Lymphocites Lysine MAC-1 Morphogens Nitricoxide (NO) Nucleotides n-methylpyrrolidinone (NMP) Dimethyl Sulfoxide(DMSO) Peptides Phosphorylcholine Phospholipids Polypeptides PR39Proteins Prostaglandins Proteoglycans Perlecan Radioactive materialsIodine - 125 Iodine - 131 Iridium - 192 Palladium 103Radio-pharmaceuticals Secondary Messengers Ceramide Somatomedins StatinsSteroids Sulfonyl Thrombin Thrombin inhibitor Thrombolytics TiclidTyrosine kinase Inhibitors ST638 AG-17 Vasodilator Histamine ForskolinNitroglycerin Vitamins E C Yeast

TABLE 2 Examples of Materials and Additives Aliphatic polyestersCellulose Bioglass Chitin Collagen Copolymers of glycolide Copolymers oflactide Elastin Fibrin Glycolide/l-lactide copolymers (PGA/PLLA)Glycolide/trimethylene carbonate copolymers (PGA/TMC) HydrogelLactide/tetramethylglycolide copolymers Lactide/trimethylene carbonatecopolymers Lactide/ε-caprolactone copolymers Lactide/σ-valerolactonecopolymers L-lactide/dl-lactide copolymers Methyl methacrylate-N-vinylpyrrolidone copolymers Modified proteins Nylon-2 Organic SolventsPHBA/γ-hydroxyvalerate copolymers (PHBA/HVA) PLA/polyethylene oxidecopolymers PLA-polyethylene oxide (PELA) Poly (amino acids) Poly(trimethylene carbonates) Polyhydroxyalkanoate polymers (PHA)Poly(alklyene oxalates) Poly(butylene diglycolate) Poly(hydroxybutyrate) (PHB) Poly(n-vinyl pyrrolidone) Poly(ortho esters)Polyalkyl-2-cyanoacrylates Polyanhydrides PolycyanoacrylatesPolydepsipeptides Polydihydropyrans Poly-dl-lactide (PDLLA)Polyesteramides Polyesters of oxalic acid Polyglycolide (PGA)Polyiminocarbonates Polylactides (PLA) Poly-l-lactide (PLLA)Polyorthoesters Poly-p-dioxanone (PDO) Polypeptides PolyphosphazenesPolysaccharides Polyurethanes (PU) Polyvinyl alcohol (PVA)Poly-β-hydroxypropionate (PHPA) Poly-β-hydroxybutyrate (PBA)Poly-σ-valerolactone Poly-β-alkanoic acids Poly-β-malic acid (PMLA)Poly-ε-caprolactone (PCL) Pseudo-Poly(Amino Acids) Starch Trimethylenecarbonate (TMC) Tyrosine based polymers Alginate Calcium CalciumPhosphate Calcium Sulfate Ceramics Chitosan Cyanoacrylate CollagenDacron Demineralized bone Elastin Keratin Plasticizers Fibrin GelatinGlass Gold Glycosaminoglycans Hyaluronic acid Hydrogels HydroxyapatiteHydroxyethyl methacrylate Hyaluronic Acid Liposomes Lipids NitinolNanoparticles Osteoblasts Oxidized regenerated cellulose Phosphateglasses Polyethylene glycol Polyester Polysaccharides Polyvinyl alcoholPlatelets, blood cells Radiopacifiers Salts Silicone Silk Steel (e.g.Stainless Steel) Synthetic polymers Thrombin Titanium Silica Clay MetalsSilver Aluminum Oxides Ceramics Polymers Metal Oxides Alkylmethlacrylates Hydrophilic polymer Integrins Paralyne PolyacrylamidePolyanhydrides Polyethylene acetate Polyethylene glycol Polyethyleneoxide Polyurethane Polyvinyl alcohol Polyvinyl pyrrolidone SilanesSilicone

1. A multiphasic device for repair or replacement of articular cartilageand the underlying bone, said device comprising a bone region and amalleable cartilage region said bone region comprising a porous materialand a rigid penetrating force conductive material for conductingpressure forces experienced by said cartilage region through said boneregion and into said external underlying bone tissue, thereby preventingbone tissue loss in said external underlying bone tissue.
 2. The deviceof claim 1 wherein the pressure forces are conducted as hydrostaticpressure pulses along rigid conductive elongated channels runningthrough the porous material of the bone phase and radiating out and awayfrom the cartilage phase.
 3. The device of claim 1 wherein the rigidconductive material is in the form of one or more vertical rigid columnsoriented perpendicular to the top and bottom surfaces that flare out asthey reach the underlying external bone, said vertical rigid columnstransforming the fluid forces into kinetic pressure pulses that arecarried through the porous material and into the underlying bone.
 4. Thedevice of claim 1 wherein the rigid conductive material is a rigidmulti-facetted web structure oriented perpendicular to the top andbottom surfaces, said web structure transferring the fluid forces intokinetic pressure pulses that are carried through the porous material andinto the underlying bone.
 5. The device of claim 1 wherein the rigidconductive material is in the form of one or more vertical rigid conesor wedges oriented perpendicular to the top and bottom surfaces with thebase located towards the cartilage phase, said vertical rigid columnstransforming the fluid forces into kinetic pressure pulses that arecarried through the porous material and into the underlying bone whileat the same time applying an outward force to the porous material. 6.The device of all the above claims wherein the rigid conductive materialextends partially into the malleable cartilage phase.
 7. The device ofall the above claims wherein the rigid conductive material extendspartially into the external underlying bone.
 8. The device of all theabove claims wherein at least one of the materials is bioresorbable 9.The device of all the above claims wherein the rigid conductive materialis porous.
 10. The device of all the above claims wherein the rigidconductive material is selected from the groups consisting of metals,ceramics, polymers, glasses, or combinations thereof.
 11. The device ofall the above claims wherein the porous material is a polymer.
 12. Thedevice of all the above claims wherein the porous material contains anadditive in the form of a particulate or biologically active agent. 13.The device of all the above claims wherein the rigid conductive materialcontains an additive in the form of a particulate or biologically activeagent.
 14. A multiphasic device for repair or replacement of articularcartilage and the underlying bone, said multiphasic device comprising abone region and a malleable cartilage region, said bone regioncomprising at least a first material in the form of at least two porousrigid scaffolds separated by at least a second material in the form of amalleable elastic hydrogel, said hydrogel being capable of transferringhydrostatic pressure pulses through the bone region, said bone regionpreventing bone voids from forming in external underlying bone tissue byconducting pressure forces experienced in said malleable cartilage phasethrough said bone region and into said underlying bone tissue.
 15. Themultiphasic device of claim 14 wherein the porous rigid materials are inthe form of disks having a thickness of 1000 microns to 4000 microns,separated by a hydrogel having a thickness of no less that 5 microns.16. The multiphasic device of claim 14 wherein the porous rigidmaterials are in the form of porous particles having an approximatediameter in the range of 1000 microns to 4000 microns, suspended in ahydrogel, and separated from each other by no less than 5 microns. 17.The multiphasic device of claim 14 wherein at least one of the materialsis bioresorbable.
 18. The multiphasic device of claim 14 wherein thedevice becomes malleable upon hydration.
 19. The device of claim 14wherein the porous rigid material is selected from the groups consistingof metals, polymers, ceramics, glasses or combinations thereof.
 20. Thedevice of claim 14 wherein the matrix contains an additive in the formof a particle or biologically active agent.
 21. The device claim 14wherein the porous rigid material contains an additive in the form of aparticulate or biologically active agent.
 22. The device claim 14wherein the hydrogel contains an additive in the form of a particulateor biologically active agent.